Blood pump flow rate control method and apparatus utilizing multiple sensors

ABSTRACT

A flow rate control method and apparatus for a blood pump is provided wherein multiple sensors are utilized to measure physical dimensions of a ventricle to calculate either a cross sectional area or a volume of the ventricle during distention and contraction. The multiple sensors are positioned at spaced locations around the ventricle, and may be attached to or imbedded in the ventricle. The sensors may be in the same plane for determining a cross sectional area or in different planes for determining a volume. The area or volume is generally proportional to a corresponding area or volume of the ventricle such that the calculated area or volume is indicative of distention and contraction of the volume. The area or volume is utilized to control the blood pump flow rate to avoid overly distending or contracting the ventricle and for operating the flow rate in a pulsatile manner to closely approximate the natural pumping action of the heart.

CROSS REFERENCE TO RELATED APPLICATION

This application is a continuation-in-part of U.S. patent applicationSer. No. 09/273,384, filed Mar. 22, 1999, and allowed on Nov. 7, 2001,which is a divisional application of U.S. patent application Ser. No.08/978,670, filed Nov. 26, 1997, now U.S. Pat. No. 5,928,131.

BACKGROUND

1. Field of the Invention

The present invention relates generally to rotary pumps utilized to pumpblood, and more particularly to a flow rate control method and apparatusutilizing multiple sensors for use in determining and area or volume ofa heart ventricle during natural distention and contraction andutilizing that information to control the flow rate of a blood pump.

2. Description of the Prior Art

The use of rotary pumps (i.e. axial, centrifugal, mixed flow) to pumpfluids and in particular blood is well known by those skilled in theart. A rotary pump, in general, consists of an outer housing, with inletand outlet ports, and an impeller mounted on a shaft (with mechanicalbearings and seals) within the outer housing for rotation about an axis.Mechanical bearings are susceptible to wear and premature failure andcan generate sufficient heat and mechanical stresses to causeunacceptable blood damage. Shaft seals are also susceptible to wear andheat generation, which can lead to leakage, blood clot formation,bearing seizure, and bacterial growth. Examples of rotary pumpsutilizing shaft mounted impellers with bearings and seals are disclosedin Reich et. al. U.S. Pat. No. 4,135,253; Possell U.S. Pat. No.4,403,911; Moise U.S. Pat. No. 4,704,121; and Dorman U.S. Pat. No.4,927,407.

Numerous pumps have been designed to circumvent the above problems byemploying a lubricant flush of rotary pump mechanical bearings. Examplesof such pumps are disclosed in Carriker et al. U.S. Pat. No. 4,944,722and Wampler et al. U.S. Pat. No. 4,846,152. These types of pumps canhave several problems including not having the ability to be fullyimplantable due to the need for a percutaneous supply line and externalreservoir to achieve bearing flushing. Also the potential for infectionand leakage exists due to the flushing fluid and percutaneous lines. Inaddition the mechanical bearings can still require replacement aftertime because they directly contact other pump structures duringoperation.

By employing a rotary fluid pump with a magnetically suspended impeller,all of the above mentioned problems can be avoided. Examples of suchpumps are disclosed in Bramm et al. U.S. Pat. No. 5,326,344; Olsen etal. U.S. Pat. No. 4,688,998 and Moise U.S. Pat. No. 4,779,614. A problemwhich can be associated with all of the cited inventions is that asingle gap is employed for both the blood flow pathway through the pumpand for the magnetic suspension and rotation of the impeller. These twofunctions have directly opposing requirements on the size of the gap. Asa blood flow pathway, the gap should be large to avoid blood damage. Asa magnetic suspension and rotation gap, the gap should be small tominimize the size of the magnetic suspension and rotation components andalso to allow for efficient use of energy to achieve impeller suspensionand rotation. Consequently, for these types of pumps, any gap sizeselected can result in an undesirable compromise between blood damage,device size, and energy requirements.

Examples of pumps having separate gaps for primary blood flow andimpeller rotation are disclosed in Golding et al. U.S. Pat. No.5,324,177 and Golding et al. U.S. Pat. No. 5,049,134. However, thesepumps also use the rotation gap to implement hydrodynamic suspensionbearings for the rotor. Such hydrodynamic bearings can subject the bloodto excessive shear stresses which can unacceptably damage the fragilecomponents of the blood. Additionally, the Golding et. al. pumps placethe stationary magnetic components inside a center-bore of a rotatingassembly. Such configurations generally cause the mass and rotationalinertia of the rotating assembly to be larger than those in a system inwhich the stationary magnetic components are placed around the outersurface of the rotating assembly. Rotating assemblies having largemasses and rotational inertias can be undesirable because the axial andradial bearing elements must be made relatively large in order tomaintain proper alignment of the rotating assembly during shock,vibration, and acceleration.

The flow rate of blood pumps that are capable of creating negative inletpressures must be dynamically adjusted to match the blood flow rate intothe ventricle of the heart, typically the left ventricle. If too littleflow is produced by the blood pump, the tissues and organs of the bodymay be inadequately perfused, and the blood pressure in the leftventricle will increase—potentially causing excessive pulmonary pressureand congestion. Conversely, if the flow rate of the blood pump is toohigh, excessive negative pressure may be created in the left ventricleand in the inlet to the pump. Excessive negative blood pressure isundesirable for the following reasons: 1) Unacceptable levels of blooddamage may be caused by cavitation; 2) The pump may be damaged bycavitation; 3) The walls of the ventricle may collapse and be damaged;and 4) The walls of the ventricle may collapse and block the blood flowpathway to the pump.

By employing a control system to dynamically control the flow rate ofthe pump to avoid excessive negative blood pressure the above mentionedproblems can be avoided. One example of such a control system isdisclosed in Bramm et al., U.S. Pat. No. 5,326,344. Bramm describes amethod of dynamically controlling the flow rate of a pump based on asignal derived from a single pressure sensor located within the pumpinlet. One problem which can be associated with such a pressure sensingsystem is the difficulty in achieving long-term stability of such asensor, particularly in light of the relatively low pressures (0 to 20mm Hg) that must be resolved and the hostile environment in which thesensor is operated. Another problem which can be associated with such apressure sensing system is that the effect of changing atmosphericpressure can cause inaccurate sensing of the pressure needed to properlycontrol the pump.

Many patients that are in need of cardiac assistance due to theirheart's inability to provide adequate blood flow are also predisposed tocardiac arrhythmias. Such arrhythmias can adversely affect blood flowwhen a cardiac assist device is used, particularly when onlyuni-ventricular cardiac assistance is being provided. By combining anarrhythmia control system with a cardiac assistance system, the abovementioned problems can be alleviated. One example of such a combinedcardiac assist and arrhythmia control system is disclosed by Heilman etal. U.S. Pat. No. 4,925,443. Heilman describes a cardiac assist devicethat directly compresses the myocardium to achieve increased blood flowcombined with an arrhythmia control system. Some problems which can beassociated with direct compression of the myocardium can includedifficulty in conforming to a wide range of heart shapes and sizes,difficulty in adequately attaching such a device to the heart, anddamage of the myocardium due to compression and abrasion.

Accordingly, there is a need for a blood pump which overcomes theaforementioned problems that can be associated with conventional bloodpumps and also a system of dynamically controlling such a blood pump toavoid the previously described problems that can occur with controlsystems using pressure sensors. Moreover, such blood pump and controlsystem should be able to cooperate with an arrhythmia control system forimproved cardiac arrhythmia treatment.

SUMMARY

A flow rate control method and apparatus for use with a blood pump isprovided wherein multiple sensors can be employed to sense physicaldimensions of a heart ventricle to calculate either a cross sectionalarea or a volume of the ventricle during distention and contraction. Themultiple sensors can be positioned at selected locations around theventricle, and may be attached to or imbedded in the ventricle. Thesensors can be ultrasonic transducers positions at selected pointsaround the left ventricle for obtaining measurements to approximate across sectional area of the ventricle. Each of the sensors can bepositioned in a plane through the ventricle where the physicalcharacteristic to be determined is a cross sectional area. Where thephysical characteristic to be determined is a volume, the sensors can bepositioned in different planes. In either case, the area or volume isgenerally proportional to a corresponding area or volume of theventricle such that the calculated area or volume is indicative ofdistention and contraction of the volume. The calculated area or volumecan be a more precise indication of distention and contraction than asingle radial dimension measurement. The area or volume can be utilizedto control the blood pump flow rate to avoid overly distending orcontracting the ventricle and for operating the flow rate in a pulsatilemanner to closely approximate the natural pumping action of the heart.

Although any type of suitably controllable blood pump could becontrolled using the method and apparatus described above, the bloodpump apparatus can preferably include a stator member containing amagnetically suspended and rotated rotor member. The rotor canpreferably be magnetically suspended within the stator both radially andaxially. The blood pump can also have an associated magnetic suspensioncontrol system, a blood pump flow rate control system, and an arrhythmiacontrol system. The blood pump can preferably be a centrifugal pumpwherein an impeller draws blood from the left ventricle of a the heartand delivers it to the aorta thereby reducing the pressure that must begenerated by the left ventricle. The blood pump can also be of arelatively small size such that it can be completely implanted withinthe human body. If bi-ventricular cardiac assist is needed a second suchblood pump can be implanted to assist the right ventricle. The impellerof the centrifugal pump can be an integral part of a rotor assembly. Therotor assembly can preferably be suspended by permanent magnet radialbearings and a Lorentz-force axial bearing. The Lorentz-force axialbearing can generate bi-directional axial forces in response to anapplied current. The blood pump can also include an axial positionsensor and an axial position controller. The axial position sensor canmonitor the axial position of the rotor and provide feedback to thecontroller to maintain the axial position of the rotor. The axialposition controller can also adjust the axial position of the rotor suchthat steady-state axial loads due to gravity, acceleration or thecentrifugal pump impeller are offset by the inherent axial forcesgenerated by the permanent magnet radial bearings. By offsetting thesteady-state axial forces using the axial position controller, the powerrequired by the Lorentz-force axial bearing is minimized. The rotorassembly can be rotated by an electric motor.

A primary blood flow inlet path can preferably be through a relativelylarge center bore provided in the rotor. A secondary blood flow inletpath can be through an annular gap which is formed between the rotor andthe stator of the pump as a result of the radial magnetic suspension. Inorder to minimize the size of the device, all of the magnetic suspensionand rotation forces can be applied across the relatively small annulargap. All blood contacting surfaces of the pump are continuously washedto avoid blood clots and protein deposition.

The speed of the centrifugal pump can be dynamically controlled to avoidexcessive negative pressure in the left ventricle using the multiplesensor control method and apparatus described above. An alternativeblood pump flow rate control system can utilize an electronic heartcaliper operatively attached to the outside surface of the heart toprovide feedback to the blood pump flow rate control system. The heartcaliper generally measures a single, generally radial dimension of theventricle which is indicative of distention and contraction of theventricle.

The blood pump flow rate control system can preferably operate in twomodes, continuous and pulsatile. In the continuous mode of operation,the pump speed can be controlled to hold the sensed or calculated leftventricle dimension, area or volume at a defined setpoint. In thepulsatile mode of operation, the pump speed can be dynamically adjustedto cause the sensed or calculated left ventricle dimension, area orvolume to alternate between two predefined setpoints.

The blood pump can also be utilized to improve the functioning of anarrhythmia control system. Electrodes placed in or on the surface of theheart combined with an associated arrhythmia control system can beprovided to detect and treat cardiac arrhythmias including bradycardia,tachycardia, and fibrillation. In order to reduce the energy needed forthe arrhythmia control system to treat fibrillation, the blood pump flowrate control system can be employed to purposely reduce the radialdimension of the ventricle prior to delivering a defibrillation pulse.By minimizing the amount of blood within the ventricle chamber (a directresult of reducing the radial dimension thereof), a larger fraction ofthe defibrillation energy supplied by the arrhythmia control system isdelivered to the myocardium, where it is needed, and a smaller fractionof the energy is delivered to the blood, where it is unnecessary.

Other details, objects, and advantages of the invention will becomeapparent from the following detailed description and the accompanyingdrawing figures of certain presently preferred embodiments thereof.

BRIEF DESCRIPTION OF THE DRAWINGS

A more complete understanding of the invention can be obtained byconsidering the following detailed description in conjunction with theaccompanying drawings, wherein:

FIG. 1 is a cross section view of an embodiment of the blood pump havinga magnetically suspended and rotated rotor assembly;

FIG. 2 is a view of the blood pump in FIG. 1 taken along line II—II;

FIG. 3 is a view of the blood pump shown in FIG. 2 having adouble-volute configuration;

FIG. 4 is a perspective view of the blood pump of FIG. 1 connected to acirculatory system;

FIG. 5 is a schematic diagram of a circuit for sensing the axialposition of the magnetically suspended rotor assembly;

FIG. 6 is a simplified schematic diagram of an axial positioncontroller;

FIG. 7 is a graphical illustration of a minimum power axial positioncontrol method;

FIG. 8a is a sectional view of a heart caliper attached to a distendedventricle;

FIG. 8b is a sectional view of a heart caliper attached to a contractedventricle;

FIG. 9 is an enlarged sectional view of an apparatus for electronicallymeasuring the angle between two caliper arms shown in FIGS. 8a and 8 b;

FIG. 10a is a sectional view of a sonomicrometry based heart caliperattached to a distended ventricle;

FIG. 10b is a sectional view of a sonomicrometry based heart caliperattached to a contracted ventricle;

FIG. 11 is a graphical illustration of a method for controlling asteady-state flow rate of the blood pump; and

FIG. 12 is a graphical illustration of a method for controlling the flowrate of the blood pump in a pulsatile manner;

FIG. 13a is a sectional view of an alternative heart measurementapparatus for determining a cross sectional area utilizing multiplesonomicrometry based sensors associated with a distended ventricle;

FIG. 13b is a sectional view of the heart measurement apparatus of FIG.13a associated with a contracted ventricle;

FIG. 14a is a sectional view of an alternative heart measurementapparatus for determining a volume utilizing multiple sonomicrometrybased sensors associated with a distended ventricle;

FIG. 14b is a sectional view of the heart measurement apparatus of FIG.14a associated with a contracted ventricle.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Referring now to the drawing figures wherein like reference numbersrefer to similar parts throughout the several views, a presentlypreferred blood pump apparatus is shown in FIG. 1 having a statorassembly 1 and a rotor assembly 2.

The stator assembly 1 can have an outer stator shell 3, an inner volutehousing 4, an outer volute housing 5, and a thin-walled stator liner 6.The stator shell 3, inner volute housing 4 and stator liner 6 can eachbe made from titanium. The stator liner 6 can have a thickness fromabout 0.005 to 0.015 inch, and preferably is about 0.010 inch. The outerstator shell 3, an inner volute housing 4, and stator liner 6 canpreferably be welded together to form a hermetically sealed annularstator chamber 54. The stationary magnetic suspension and motorcomponents can be advantageously housed in the stator chamber 54.

The rotor assembly 2 can have a relatively large center bore 22 whichcan be the primary blood flow path 22 through the pump. Preferably thecenter bore 22 is about 0.50 inch. The rotor assembly 2 can include aninner rotor support sleeve 7, a rotor end cap 8 and a thin-walled rotorliner 9. The inner rotor support sleeve 7, rotor end cap 8 and rotorliner 9 can each be made from titanium. The rotor liner 9 can have athickness from about 0.005 to 0.015 inch and can preferably be about0.010 inch. The rotor support sleeve 7, rotor end cap 8 and rotor linercan preferably be welded together to form a hermetically sealed annularrotor chamber 55. The rotating magnetic suspension and motor componentscan be advantageously housed in the rotor chamber 55. The inner rotorsupport sleeve 7 can be fabricated with an integral impeller 10 or,alternately, the impeller 10 can be fabricated independently then weldedor bonded to the rotor support sleeve 7.

The blood contacting surfaces of the blood pump can be coated with adiamond-like carbon film or a ceramic film. Such films enhance the longterm bio-compatibility of the surfaces by improving their surface finishand abrasion resistance. Companies capable of providing such filmsinclude Diamonex Performance Products, Allentown, Pa, and ImplantSciences Corporation, Wakefield, Mass.

The primary inlet blood flow path 20, can be through the center bore 20of the inner rotor support sleeve 7. A secondary inlet blood flow path21, can be through the annular gap 21 which is the radial magneticsuspension gap between the stator liner 6 and the rotor liner 7. Theannular gap 21 can preferably be about 0.020 inch. The blades of theimpeller 10 can include outer portions 52, that purposely draw bloodthrough the secondary inlet blood flow path 21, and inner portions 53,that purposely draw blood through the primary inlet blood flow path 20.

When polarized as indicated in FIG. 1, radial magnetic repulsion forcesare generated between permanent magnets 11, mounted in the statorchamber 54, and permanent magnets 12, mounted in the rotor chamber 55.As the rotor assembly 2, is moved radially downward relative to thestator assembly 1, the repulsion force between the lower portion 29, ofpermanent magnets 11 and 12 increases while the repulsion force betweenthe upper portion 30, of permanent magnets 11 and 12 decreases. A netupward force is thus created which tends to restore the rotor assembly2, to a radially aligned position. Likewise, as the rotor assembly 2, ismoved radially upward relative to the stator assembly 1, the repulsionforce between the upper portion 30, of permanent magnets 11 and 12increases while the repulsion force between the lower portion 29, ofpermanent magnets 11 and 12 decreases. A net downward force is thuscreated that tends to restore the rotor assembly 2 to the radiallyaligned position. The described radial repulsion forces tend to causethe rotor assembly 2 to remain radially suspended with respect to thestator assembly 1. Permanent magnets 11 and 12 can preferably befabricated of magnetically hard material having a relatively high energyproduct such as Neodymium Iron Boron.

An assembly of permanent magnets 13, 14, coils 16, 17, and back irons15, 18 cooperate to form a Lorentz-force actuator which can be employedas an axial bearing to support the rotor assembly 2 axially. Permanentmagnets 13 and 14 cause magnetic flux 19, to flow radially from theouter surface of magnet 13, radially across the secondary blood flowpath 21, radially through coil 16, axially through the stationaryactuator back-iron 18, radially through coil 17, radially across thesecondary blood flow path 21, radially through magnet 14, axiallythrough the rotating actuator back-iron 15, and radially through magnet13. Permanent magnets 13 and 14 can preferably be fabricated of amagnetically hard material having a relatively high maximum energyproduct such as Neodymium Iron Boron and can preferably be bonded to therotating actuator back-iron 15, which in turn can preferably be bondedto the inner rotor support sleeve 7. The stationary actuator back-iron18, and rotating actuator back-iron 15, can preferably be fabricated ofa soft magnetic material having a high saturation flux density. One suchmaterial is 48% Iron-48% Cobalt-2% Vanadium available as HIPERCO® 50Afrom Carpenter Technology Corporation, Reading Pa. Coils 16 and 17 canbe fabricated from copper or silver wire and can preferably be bonded tothe stationary actuator back-iron 18, which in turn can be bonded to theouter stator shell 3. When coils 16 and 17 are energizing such thatcurrent flows in a clockwise direction in coil 16 and in acounterclockwise direction in coil 17, as viewed from the pump inlet 22,a net axial Lorentz force is generated which tends to move the rotorassembly 2 to the right. If the direction of the currents in coils 16and 17 is reversed such that current flows in a counterclockwisedirection in coil 16 and in a clockwise direction in coil 17, as viewedfrom the pump inlet 22, a net axial Lorentz force is generated whichtends to move the rotor assembly 2, in the left. A Lorentz-forceactuator as described can be preferable to attractive ferromagneticactuators because: a single Lorentz-force actuator is capable ofproducing bi-directional forces; the force output is a linear functionof input current; the bandwidth is wider; the attractive radial forcebetween the moving and stationary portions of the actuator is relativelylow; and the generated force is parallel to the axial gap formed betweenthe moving and stationary portions of the actuator.

A permanent magnet 31, armature windings 32 and back-iron 33 cooperateto form a slotless, brushless DC motor with a coreless armature. Suchslotless, coreless motors are well understood by those skilled in theart and are described in U.S. Pat. No. 4,130,769. A 2-pole permanentmagnet ring 31 causes magnetic flux to flow radially from the its northpole 34, across the secondary blood flow gap 21, radially through thearmature windings 32, circumferentially through the stator back-iron 33,radially through the armature windings 32, radially across the secondaryblood flow gap 21 to the south pole 35 of the permanent magnet ring 31.Interaction between axial current flowing in the armature windings 32and the radial magnetic flux produces torque between the rotor assembly2 and the stator assembly 1. The permanent magnet ring 31 can preferablybe fabricated of a magnetically hard material having a relatively highmaximum energy product such as Neodymium Iron Boron. Alternatively, thepermanent magnet ring 31 can be replaced with permanent magnet ringassemblies having more than 2 poles in order to reduce the size and/orincrease the efficiency of the motor. The stator back-iron assembly 33,can be fabricated from a stack of magnetically soft lamination ringspreferably having high resistivity and a high saturation flux density.One such material is 48% Iron-48% Cobalt2% Vanadium and is available asHIPERCO® 50A from Carpenter Technology Corporation, Reading Pa.Electrically insulated laminations are used in the stator back-ironassembly 33 to minimize power losses caused by eddy currents which areinduced by the rotating magnetic field produced by permanent magnet ring31. It is understood that a conventional salient-pole brushless DC motorcould be used in place of the described motor, however, a slotless,coreless, motor can be preferable because cogging torque can beeliminated in slotless motors allowing smoother, quieter operation ascompared to salient-pole brushless DC motors, and slotless, coreless,motors generally have larger radial gaps between the permanent magnetsin the rotor and the stator back-iron resulting in lower attractiveradial forces. Attractive radial forces generated by the motor can beundesirable since they tend to oppose the repulsive radial suspensionforces generated by the permanent magnet radial bearing magnetsresulting in reduced radial suspension stiffness. Such slotless,brushless, coreless DC motors are available from companies such asElectric Indicator Company, Inc. Norwalk, Conn.; Portescap U.S., Inc.,Hauppauge, N.Y.; Maxon Precision Motors, Inc., Fall River, Mass.; andMicroMo Electronics, Inc., Clearwater, Fla.

An assembly of coils 23, 24 and ferromagnetic rings 25, 26 cooperate toform an axial position sensor which is used to monitor the axialposition of the rotor assembly 2 with respect to the stator assembly 1.The two coils 23, 24 can be fabricated from copper wire. A firstferromagnetic ring 25 causes the inductance of a first coil 23 toincrease and the inductance of the second coil 24 to decrease as it ismoved to the left. Likewise, the inductance of the first coil 23decreases and the inductance of the second coil 24 increases as thefirst ferromagnetic ring 25 is moved to the right. A secondferromagnetic ring 26 can serve to both magnetically shield and increasethe Q of the coils 23, 24. The two ferromagnetic rings 25, 26 canpreferable be made of a ferrite material having a high permeability atthe excitation frequency of the coils 23, 24. One such material isMATERIAL-W® available from Magnetics, Division of Spang & Co., Butler,Pa. A pair of spacers 27, 28 can be used to radially locate the twoferromagnetic rings 25, 26.

An annular pump chamber having a single volute passage, is shown in FIG.2. The annular pump chamber is shown having an outer volute housing 5and an inner volute housing 4. The impeller 10 rotates within the pumpchamber about a projection of the outer volute housing 5 and within theinner volute housing 4. A series of impeller blades 36 propel blood fromthe primary blood flow path 20, centrifugally around the volute passage37, and out the outflow port 38.

The single-volute centrifugal pump illustrated in FIG. 2 inherentlydevelops a radial force on the impeller which must be offset by thepermanent magnet radial bearings 11, 12, shown in FIG. 1. To minimizethis radial force, an alternative, double volute, configuration, asshown in FIG. 3, can be employed. A double volute passage can be formedin the annular pump chamber by interposing a septum 56 in the singlevolute passage 37 (shown in FIG. 2) to form a pair of volute passages39, 40 which are radially opposed. It is to be understood that theoverall size of the double volute passages and outlet 42 may be largerthan single volute passage 37 and outlet 41 to accommodate the septum 56and allow for adequate blood flow through the annular pump chamber. Theradially-opposed volutes 39, 40 produce opposing impeller forces thatbalance one another and thus minimize the radial force that must beoffset by the permanent magnet radial bearings 11, 12. In the doublevolute configuration, similarly to the single volute design, theimpeller 10 rotates about a projection of the outer volute housing 5 andwithin the inner volute housing 4. However, the impeller blades 36 nowpropel blood from the primary blood flow path 20 through bothcentrifugal volute passages 39, 40. The blood, flowing separately ineach volute passage 39, 40, combines at confluence point 41 and isdelivered to the outlet 42. It should be understood that otherimpeller-volute arrangements could be derived by those skilled in theart and the invention is not to be limited to the particularconfigurations illustrated and described herein.

Referring now to FIG. 4, one method of connecting of the blood pump 51to the circulatory system is schematically illustrated. Several cannulas44, 46, 49 can be provided to connect the pump 51 between the leftventricle of the heart and the aorta. A hole is cored in the apex of theleft ventricle at location 43 and one cannula 44 directs blood from theleft ventricular cavity to the pump inlet 45. Another cannula 46 directsblood from the pump outlet 47 to an in-line artificial heart valveassembly 48. Alternatively, a solenoid actuated valve could be used inplace of valve assembly 48. The artificial heart valve assembly 48 canpreferably be provided to prevent retrograde blood flow from the aorta,through the pump, and into the left ventricle in the event of a failureof the blood pump or an associated control system. From the outlet ofthe heart valve assembly 48, another cannula 49 directs the blood to theascending aorta 50. For bi-ventricular cardiac assist, a second pumpcould be connected in like fashion between the right ventricle andpulmonary artery.

An axial position sense subsystem 141 can have the circuitry shown inFIG. 5. The subsystem 141 can utilize the ratio of the inductances ofcoils 23 and 24 to measure the axial position of the rotor assembly 2.The subsystem 141 can include an amplitude-stabilized sine-waveoscillator 100, which is used to excite coils 23 and 24 arranged as ahalf-bridge 101, and a synchronous demodulator 102. Synchronousdemodulation is used to detect the relatively low amplitude signalsoutput from the half-bridge circuit 101 because the synchronousdemodulation technique effectively filters electrical noise at allfrequencies except those centered about the excitation frequency. Anoscillator 103 generates a square wave output 104, which is used tocontrol analog switch 105. The output 106, of analog switch 105 is asquare wave that alternates between the output voltage 107, ofoperational amplifier 108, and ground. A capacitor 109 and a resistor110 form a highpass filter that removes the DC offset from signal 106. Alowpass filter 111 attenuates the upper harmonics of input signal 112resulting in a sine-wave output signal 113. The lowpass filter 111 is ofsufficient order and type to attenuate the third harmonic of the squarewave input 112 by 40 dB or more. One possible configuration for low passfilter 111 is a 5^(th) order Butterworth type. A capacitor 114 removesany DC offset from the output 113 of the lowpass filter 111. An ACsine-wave 115 is used to excite the half-bridge network 101. A pair ofresistors 116, 117 and operational amplifier 1 18 form an invertingcircuit with a gain of−1. A comparator 119 detects the sign of thesine-wave excitation signal 115. The output 120 of the comparator 119 isused to control an analog switch 121. When the sign of sine-wave 115 isnegative, the output 122 of the analog switch 121 is connected to thenon-inverted sine-wave signal 115. When the sign of sine-wave 115 ispositive, the output 122, of the analog switch 121 is connected to theinverted sine-wave signal 123. The output 122 is thus the inverted,full-wave rectified representation of the excitation sine-wave signal115. An operational amplifier 108, a pair of resistors 123,124 and acapacitor 125 form an integrating difference amplifier. The output 107of the operational amplifier 108 increases if the average full-waverectified representation of the excitation sine-wave signal 115 is lessthan the applied precision reference voltage 126. Likewise the output107 of the operational amplifier 108 decreases if the average full-waverectified representation of the excitation sine-wave signal 115 isgreater than the applied precision reference voltage 126. Through thedescribed integrating action, the amplitude of the AC signal 106 iscontrolled as required to maintain the average full-wave rectifiedrepresentation of the excitation sine-wave signal 115 equal to theapplied precision reference voltage 126. As previously described, theratio of the inductances of coils 23 and 24 is a function of the axialposition of the rotor assembly 2 shown in FIG. 1. The amplitude of theoutput signal 127 of the half bridge circuit 101 formed by coils 23 and24 thus varies with the axial position of the rotor assembly 2. A pairof resistors 128, 129 and an operational amplifier 130 form an invertingcircuit with a gain of −1. The output 120, of the comparator 119 is usedto control an analog switch 131. When the sign of sine-wave 115 isnegative, the output 132 of the analog switch 131 is connected to thenon-inverted output signal 127 of the half bridge circuit 101. When thesign of sine-wave 115 is positive, the output 132 of the analog switch131 is connected to the inverted output signal 133 of the half bridgecircuit 101. The output signal 132 is thus the inverted, full-waverectified representation of the output signal 127 of the half bridgecircuit 101. A lowpass filter 134 attenuates the AC components of theoutput signal 132. One possible configuration for the low pass filter134 is an 8^(th) order Butterworth type. Several resistors 135, 136,137, along with an operational amplifier 138 and a precision referencevoltage 126 shift and scale the output 139 of the lowpass filter 134 asrequired for downstream circuits. The output 140 of operationalamplifier 138 is thus a representation of the axial position of therotor assembly 2. Consequently, changes in the output 140 provide ameasurement of the axial movement of the rotor assembly 2. The circuitillustrated in FIG. 5 is but one example of a circuit that can be usedto detect changes in the ratio of the inductances of coils 23 and 24. Itshould be understood that other acceptable circuits may be derived bythose skilled in the art.

Using the output 140 from the axial position sense subsystem 141, anaxial position controller 200, shown in FIG. 6, can be used to bothmaintain the axial position of the rotor at a defined axial positionsetpoint and to adjust the axial position setpoint to minimize powerdissipation in the Lorentz force actuator. The axial position controller200 can have the basic circuitry shown in FIG. 6, including circuitry201, which maintains the rotor at a defined axial position setpoint andcircuitry 202, which adjusts the axial position setpoint for minimumpower dissipation in the Lorentz-force actuator coils 16, 17. The axialposition setpoint maintenance circuit 201, is comprised of thepreviously described axial position sense subsystem 141, a gain andservo compensation circuit 203, a switching power amplifier 204, theLorentz-force actuator 205, and the rotor assembly 2. The axial positionsense subsystem 141 outputs a signal 140, proportional to the axialposition 215 of the rotor assembly 2. Several resistors 207, 208, 209along with a capacitor 210 and an operational amplifier 211 form a gainand lead compensation network 203, which modifies the gain and phase ofsignal 140 as required to prevent unstable oscillation of the rotorassembly 2. The design of such gain and lead compensation networks iswell understood by those skilled in the art of servo system design. Thevoltage output 212 of the gain and lead compensation network 203 isinput to switching power amplifier 204. Switching power amplifier 204outputs a current signal 213 that is proportional to the input voltage212. The design of such transconductance switching amplifiers is wellunderstood by those skilled in the art. The current signal 213 isapplied to the coils 16, 17 of the Lorentz-force actuator 205. TheLorentz-force actuator 205 produces an axial force 214 proportional tothe applied current signal 213. The axial force 214 is applied to therotor assembly 2. The axial position 215 of the rotor assembly 2 changesin response to the applied axial force 214. The overall polarity of thedescribed servo loop 201 is such that the force produced by theLorentz-force actuator opposes displacement of the rotor assembly fromthe defined setpoint. Those skilled in the art will recognize that thefunction of the analog, gain and servo compensation circuit 203 can beimplemented with software running on a microprocessor or digital signalprocessor.

In FIG. 7, the described minimum axial control power method isillustrated. The x-axis 300 of the graph represents the axial positionof the rotor assembly 2 relative to the stator assembly 1, as shown inFIG. 1. The y-axis 301 of the graph represents the axial force appliedto the rotor assembly 2. Line 302 represents the inherent axial forcesgenerated by the permanent magnets 11, 12 for small axial displacementsof the rotor assembly 2. At point 303 on the graph, the permanentmagnets 11, 12 are magnetically aligned and generate no axial force. Theslope of curve 302 is dependent on the design of the permanent magnets11, 12 and may be between 0.2 lb/0.001 inch to 1.0 lb/0.001 inch. Line304 of FIG. 7 represents a steady-state axial load applied to the rotorassembly 2. The steady-state axial load 304 may be caused by gravity,acceleration, the centrifugal pump impeller, etc. Line 305 of FIG. 7 isthe addition of lines 302 and 304 and represents the net force versusaxial position of the rotor assembly 2 when the steady-state load 304 isapplied. Point 306 defines the axial position of the rotor assemblywhere the steady-state load force is canceled by the axial forceproduced by the permanent magnets 11,12. By adjusting the axial positionsetpoint of the rotor assembly 2 to the axial position defined by point306, the steady-state actuator force output required to maintain theaxial setpoint is zero. Since the power dissipated by the Lorentz-forceactuator is proportional to the square of its output force, the netpower dissipated by the actuator is minimized when the rotor assembly isoperated at the axial position defined by point 306. Likewise, with nosteady state load forces applied, the net power dissipated by theactuator is minimized when the rotor assembly is operated at the axialposition defined by point 303.

The circuitry 202, shown in FIG. 6, can be employed to effectivelyadjust the axial setpoint position of the rotor assembly 2 for minimumpower dissipation in the Lorentz-force actuator 205 using the previouslydescribed method. The steady-state axial position setpoint can becontrolled by the voltage output 216 of the operational amplifier 217and the resistor 218. The circuit formed by the resistor 219, capacitor220 and the operational amplifier 217 inverts and integrates the voltageoutput 212 of the gain and lead compensation network 203. Signal 212 isdirectly proportional to the current flowing in the Lorentz-forceactuator coils 16, 17. If the average voltage of signal 212 is positive,indicating that a net positive current is flowing in the actuator coils16, 17, the output 216 of the operational amplifier 217 decreases andshifts the axial setpoint position of the rotor assembly 2 until theaverage current flowing in the actuator coils 16, 17 is zero. Likewise,if the average voltage of signal 212 is negative, indicating that a netnegative current is flowing in the actuator coils 16,17, the output 216of the operational amplifier 217 increases and shifts the axial setpointposition of the rotor assembly 2 until the average current flowing inthe actuator coils 16,17 is zero. The steady-state axial setpointposition of the rotor assembly 2 is thus adjusted as required forminimum power dissipation in the Lorentz-force actuator 205. Thoseskilled in the art will recognize that the function of the analog,automatic setpoint adjustment circuitry 202 can be implemented withsoftware running on a microprocessor or digital signal processor.

The flow rate of any blood pump that is capable of creating negativeinlet pressures must be dynamically adjusted to match the blood flowrate into the left ventricle. If too little flow is produced by theblood pump, the tissues and organs of the body may be inadequatelyperfused, and the blood pressure in the left ventricle will increasepotentially causing excessive pulmonary pressure and congestion.Conversely, if the flow rate of the blood pump is too high, excessivenegative pressure may be created in the left ventricle and in the inletto the pump. Excessive negative blood pressure is undesirable for thefollowing reasons: 1) Unacceptable levels of blood damage may be causedby cavitation, 2) The pump may be damaged by cavitation, 3) The walls ofthe ventricle may collapse and be damaged, and 4) The walls of theventricle may collapse and block the blood flow pathway to the pump.Preferably, the flow rate of the blood pump can be dynamicallycontrolled to avoid these problems.

A pump flow rate controller for the blood pump can be provided tooperate the pump such that the flow rate does not overly distend orcontract the ventricle. Preferably, a heart measurement apparatus canprovide the flow rate controller with information about the dimension ofthe ventricle during normal distention and contraction. Such a heartmeasurement apparatus can be an electronic heart caliper, two types ofwhich are illustrated in FIGS. 8a- 10 b.

In FIG. 8a, a cross section of a heart is illustrated, including a rightventricle 405 and a left ventricle 404 that is maximally distended bythe pressure of the blood contained therein. In FIG. 8b, the leftventricle 404 has been partially depressurized. As blood is withdrawnfrom the left ventricle 404 the radial dimension of the outside surface418 of the heart is reduced. By dynamically adjusting the flow rate ofthe blood pump to avoid excessive distention or contraction of the leftventricle, as indicated by the radial dimension of the exterior surfaceof the left ventricle, the average blood pump flow rate can becontrolled to match the flow rate of blood into the left ventricle. Oneembodiment of an electronic heart caliper 400 is shown which can beemployed to measure the radial dimension of the outside surface 418 ofthe heart. The heart caliper 400 can include two arms 401, 402 that canbe suitably attached to the outside surface 418 of the heart and pivotabout a point which can preferably be located inside a hermeticallysealed enclosure 403. A measure of the radial dimension of the leftventricle 404 can be achieved by electronically measuring the anglebetween the caliper arms 401, 402. An angular measurement apparatuswhich can be used to measure the angle between the caliper arms 401, 402is illustrated in FIG. 9. The angle measuring apparatus can preferablybe contained within a hermetically sealed enclosure 403 in order toprotect the internal components from the tissues and fluids of the body.A bellows 407, and end caps 408, 409 can preferably be welded togetherto form a hermetically sealed chamber. The bellows 407, and end caps408, 409 are preferably made from titanium. A hermetic electricalfeedthrough 410, which can use either a glass or brazed ceramicinsulator 411, can be installed in titanium end cap 409. The caliperarms 401, 402 can be effectively connected to a pivot member 415 throughrespective end caps 408, 409 and respective control arms 414, 416. Thecaliper arm 401, end cap 408 and control arm 416 can be machined from asingle piece of titanium or can be constructed individually and weldedor bonded together. Likewise, the caliper arm 402, end cap 409 andcontrol arm 414 can be machined from a single piece of titanium or canbe constructed individually and welded or bonded together. A pivot 415limits the motion of the caliper arms 401, 402 to an arc within a singleplane. As the caliper arms 401, 402 move due to distention orcontraction of the left ventricle, the extension of control arm 416moves closer or farther respectively from an eddy-current based positionsensor 417 that can be bonded to the control arm 414. The eddy-currentbased position sensor 417 can be fabricated from a miniature ferrite potcore 412 and a copper coil 413. Such miniature ferrite pot cores areavailable from Siemens Components, Inc., Iselin, N.J. The eddy-currentsensor coil can be connected to two electrical feedthroughs (only one ofthe two feedthroughs, 410, is shown in FIG. 9). As the metallic controlarm 416 moves closer to the eddy-current based position sensor 417, theeffective resistive loading of the coil increases causing a reduction ofthe coil's Q (Q is defined in the art as the ratio of reactance to theeffective series resistance of a coil). An electronic circuit can beused to measure the change in the Q of the coil and provide a signalthat corresponds to the relative position of the caliper arms 401, 402.Such electronic circuits, as described for measuring changes in Q, arewell known in the art.

An alternative embodiment of an electronic heart caliper 500 isillustrated in FIGS. 10a and 10 b. Similarly to FIGS. 8a and 8 b, FIG.10a depicts a left ventricle that is maximally distended by the pressureof the blood contained within it and Figure 10b depicts a left ventriclethat has been partially depressurized. The heart caliper 500 can have apair of arms 501, 502 that can be suitably attached to the outsidesurface of the heart and pivot about a point which can preferably belocated inside a hermetically sealed enclosure 503. A measure of theradial dimension of the left ventricle can be achieved by measuring thetime it takes for an ultrasonic pulse to travel from a sonomicrometertransducer 504 on one caliper arm 501 to an opposing sonomicrometertransducer 505 on the other caliper arm 502. Sonomicrometer transducerssuitable for use in such a heart caliper are available from companiessuch as Triton Technology, Inc., San Diego Calif. and Etalon, Inc.,Lebanon Ind. The details of sonomicrometry are well known by thoseskilled in the art. It should be understood that other suitable methodsfor measuring the relative distention and contraction of the ventricles,such as impedance and conductance measurement of the ventricle, could bederived by those skilled in the art and the invention is not to belimited to the particular methods described.

Another heart measurement apparatus that can provide the flow ratecontroller with information about the dimension of the ventricle duringnormal distention is illustrated in FIGS. 13a and 13 b. Similarly toFIGS. 8a and 8 b, FIG. 13a depicts a left ventricle 703 that ismaximally distended by the pressure of the blood contained within it andFIG. 13b depicts a left ventricle that has been partially depressurized.Three sonomicrometry transducers 700, 701, and 702 can be placed withinthe wall of the left ventricle 703. A measure of the distance 704between transducers 700 and 701 can be achieved by measuring the time ittakes for an ultrasonic pulse to travel between sonomicrometertransducers 700 and 701. A measure of the distance 705 betweentransducers 701 and 702 can be achieved by measuring the time it takesfor an ultrasonic pulse to travel between sonomicrometer transducers 701and 702. Likewise, a measure of the distance 706 between transducers 700and 702 can be achieved by measuring the time it takes for an ultrasonicpulse to travel between sonomicrometer transducers 700 and 702.Sonomicrometer transducers suitable for use in such an apparatus areavailable from companies such as Triton Technology, Inc., San Diego,Calif. and Sonometrics Corporation, London, Ontario, Canada. The detailsof sonomicrometry are well known by those skilled in the art. A measure,approximately proportional to the cross sectional area of the leftventricle can be achieved by using the measured distances 704, 705, and706 to calculate the area of the triangle 707 having vertices defined bysonomicrometry transducers 700, 701, and 702. It is preferable to placesonomicrometry transducers 700, 701, and 702 within the wall of the leftventricle 703 to most accurately approximate the cross sectional area ofthe left ventricular chamber while avoiding direct blood contact,however it is understood that ultrasonic transducers could also beplaced on the epicardial surface of the heart or in the left ventricularchamber.

Another heart measurement apparatus that can provide the flow ratecontroller with information about the dimensions of the ventricle 703during distention and contraction is illustrated in FIGS. 14a and 14 b.Similarly to FIGS. 13a and 13 b, FIG. 14a depicts a left ventricle 703that is maximally distended by the pressure of the blood containedwithin it and FIG. 14b depicts a left ventricle that has been partiallydepressurized. In addition to previously described sonomicrometrytransducers 700, 701, and 702, a fourth sonomicrometry transducer 708can be placed within the wall of the left ventricle near the apex. Ameasure of the distances 709, 710 and 711 between transducer 708 andtransducers 700, 701 and 702 respectively can be achieved by measuringthe time it takes for an ultrasonic pulse to travel betweensonomicrometer transducers. A measure, approximately proportional to thevolume of the left ventricular cavity, can be achieved by usingdistances 709, 710 and 711 along with previously described distances704, 705 and 706 to calculate the volume of tetrahedron 712 havingvertices defined by sonomicrometry transducers 700, 701, 702 and 708. Itis understood that the measure of the volume of the left ventricularcavity could be made more accurate by placing additional transducersthroughout the myocardium and calculating the volume of the definedpolyhedron.

One way to implement the flow rate controller, to control the flow rateof the blood pump to avoid excessive distention or contraction of theleft ventricle, is graphically illustrated in FIG. 11. The x-axis' 600represent time. Line 601 represents predefined distention andcontraction characteristics of the left ventricle. The predefineddistention and contraction characteristics can correspond to a radialdimension, cross sectional area or volume setpoint that is defined whenthe system is initially implanted and which may be periodically updatednoninvasively using ultrasound imaging. Curve 602 can be representativeof either the radial dimension, cross sectional area or volume of theleft ventricle as sensed by either of the previously describedelectronic heart calipers 400, 500 or the sonomicrometry systemillustrated in FIGS. 13a and 13 b. Curve 603 represents the angularvelocity of the disclosed centrifugal pump impeller. The flow rate ofthe disclosed centrifugal pump varies with the angular velocity of itsimpeller. When the sensed radial dimension, cross sectional area orvolume exceeds the corresponding setpoint as illustrated by point 604,the angular velocity of the centrifugal pump can be increased asillustrated by point 605. The increased angular velocity of thecentrifugal pump causes its flow rate to increase and more rapidlyremove blood from the left ventricle, which in turn causes the radialdimension, cross sectional area or volume of the left ventricle to bereduced towards the setpoint line 601. Likewise, when the sensed radialdimension, cross sectional area or volume is less than the setpoint asillustrated by point 606, the angular velocity of the centrifugal pumpcan be decreased as illustrated by point 607. The decreased angularvelocity of the centrifugal pump causes its flow rate to decrease andreduce the rate at which blood is removed from the left ventricle, whichin turn causes the radial dimension, cross sectional area or volume ofthe left ventricle to increase towards the setpoint line 601.

Another way to implement the flow rate controller, to control the flowrate of the blood pump to avoid excessive distention or contraction ofthe left ventricle, and also to create pulsatile blood flow, isgraphically illustrated in FIG. 12. A pulsatile flow rate more closelymimics the blood flow characteristics of a natural heart. In FIG. 12,the x-axis' 608 represent time. Lines 609 and 610 respectively representupper and lower left ventricular dimensional setpoints that are definedwhen the system is initially implanted and which may be periodicallyupdated noninvasively using ultrasound imaging. Curve 611 can berepresentative of either the measured dimension, a cross sectional areaor a volume of the left ventricle as sensed by either of the previouslydescribed electronic heart calipers 400, 500 or the multiple sensorapparatus illustrated in FIGS. 13a- 13 b. Curve 612 represents theangular velocity of the disclosed centrifugal pump impeller. The angularvelocity of the centrifugal pump can be periodically increased asindicated during time period 613. The increased angular velocity of thepump during time period 613 causes blood to be more rapidly removed fromthe heart, which in turn causes the measured dimension, cross sectionalarea or volume of the left ventricle to be reduced towards the lowerdimension setpoint line 610. The angular velocity of the centrifugalpump can be reduced as indicated during time period 614 once the senseddimension, cross sectional area or volume of the left ventricle nearlyequals the lower setpoint line 610. The reduced angular velocity of thepump during time period 615 causes the rate at which blood is removedfrom the left ventricle to be reduced, which in turn causes the senseddimension, cross sectional area or volume of the left ventricle toincrease towards the upper setpoint line 609. The angular velocity ofthe centrifugal pump can again be increased as indicated during timeperiod 616 once the sensed dimension, cross sectional area or volume ofthe left ventricle nearly equals the upper setpoint line 609. Thedescribed pulsatile pump flow rate serves to mimic the blood flowcharacteristics of the natural heart.

As part of an apparatus to automatically detect and treat cardiacarrhythmias, electrodes can be placed in or on the surface of the heartand are operatively associated with the control circuitry of theapparatus. The design and operation of such automatic arrhythmiadetection and treatment systems are well known in the art. In order toreduce the energy needed by such apparatus to treat fibrillation, theradial dimension of the ventricle can be purposely reduced just prior tothe apparatus delivering a defibrillation pulse. The radial dimension,area or volume of the ventricle can be reduced by controlling the flowrate of the previously described blood pump such that the volume ofblood within the ventricular chamber is minimized. By minimizing thevolume of blood within the ventricular chamber prior to delivering adefibrillation energy pulse, a larger fraction of the supplieddefibrillation energy can be delivered to the myocardium, where it isneeded, and a smaller fraction of the supplied defibrillation energy isdelivered to the blood, where it is unnecessary.

Although certain embodiments of the invention have been described indetail, it will be appreciated by those skilled in the art that variousmodification to those details could be developed in light of the overallteaching of the disclosure. In particular, it is to be understood thatthe blood pump flow rate control method and apparatus described hereincan be equally functional with blood pumps other than magneticallysuspended and operated blood pumps. For example, the ultrasonic heartmeasurement method and apparatus described hereinabove can be utilizedto control the flow rate of any suitably controllable blood pump,notwithstanding the particular type of pump or manner of operation.Accordingly, the particular embodiments disclosed herein are intended tobe illustrative only and not limiting to the scope of the inventionwhich should be awarded the full breadth of the following claims and anyand all embodiments thereof.

What is claimed is:
 1. An method of controlling a blood pump flow ratecomprising: a. measuring physical dimensions of a ventricle during atleast one of distention and contraction thereof; b. calculating at leastone of an area and a volume using said physical dimensions, said atleast one of an area and a volume generally proportional to acorresponding area or volume of said ventricle; c. said at least one ofan area and a volume indicative of said at least one of distention andcontraction of said ventricle; and d. controlling said blood pump flowrate based on at least one of said area and said volume.
 2. The methodof claim 1 further comprising controlling said blood pump flow rate toproduce a pulsatile flow rate.
 3. The method of claim 1 wherein saidmeasuring further comprises ultrasonically measuring a plurality ofdimensions of said ventricle.
 4. The method of claim 3 wherein saidplurality of dimensions further comprises at least three dimensionsmeasured in a single plane through said ventricle such that a crosssectional area is calculated.
 5. The method of claim 3 wherein saidplurality of dimensions are measured in more than one plane through saidventricle such that a volume is calculated.
 6. An apparatus foroperating a blood pump at a controlled flow rate comprising: a. a flowrate controller controlling said blood pump to pump blood at a flowrate; b. a heart measurement member operatively associated with saidflow rate controller and said heart measurement member attachable to aventricle of a heart which is assisted by said blood pump; c. said heartmeasurement member determining at least one of an area and a volumeduring at least one of distention and contraction of said ventricle,said at least one of an area and a volume being proportional to acorresponding area or volume of said ventricle during said at least oneof distention and contraction; and d. said flow rate controllerutilizing at least one of said area and said volume to control said flowrate.
 7. The apparatus of claim 6 wherein said flow rate is a pulsatileflow rate.
 8. The apparatus of claim 6 wherein said heart measurementmember comprises: a. a plurality of ultrasonic transducers positioned atradially spaced apart locations on said ventricle; and b. said heartmeasurement member measuring distances between each of said plurality ofultrasonic transducers, said distances utilized to determine at leastone of said area and said volume, said distances based upon an elapsedtime for an ultrasonic pulse to travel between respective ones of saidplurality of ultrasonic transducers.
 9. The apparatus of claim 8 whereinsaid plurality of ultrasonic transducers further comprises at leastthree ultrasonic transducers positioned in a single plane through saidventricle such that a cross sectional area may be determined.
 10. Theapparatus of claim 8 wherein said plurality of ultrasonic transducersare positioned in at least two planes through said ventricle such that avolume may be determined.